Two-dimensional or three-dimensional image data is often generated with the aid of advanced imaging methods, said data being used to visualize an examination object that is depicted and also for further applications.
The imaging methods are often based on the capture of X-radiation, wherein so-called projection measurement data is generated. Projection measurement data can be acquired with the aid of a computer tomography system (CT system), for example. In the case of CT systems, a combination of X-ray source and X-ray detector arranged opposite each other on a gantry usually rotate around a measurement chamber in which the examination object (subsequently referred to as the patient without thereby restricting general applicability) is situated. In this case, the center of rotation (also referred to as the “isocenter”) coincides with a so-called system axis z. During the course of one or more rotations, the patient is penetrated by X-radiation from the X-ray source, projection measurement data or X-ray projection data being captured by the X-ray detector situated opposite.
The generated projection measurement data is dependent on the design format of the X-ray detector in particular. X-ray detectors usually have a plurality of detection units which are generally arranged in the form of a regular pixel array. The detection units each generate a detection signal for X-radiation which strikes the detection units, said X-radiation being analyzed at specific time points in respect of intensity and spectral distribution in order to draw conclusions about the examination object and generate projection measurement data.
For the purpose of detecting the X-radiation, use can be made of so-called quantum counting detectors, for example. In the case of quantum counting or photon counting X-ray detectors, the detection signal for X-radiation is analyzed in respect of the intensity and the spectral distribution of the X-radiation in the form of count rates. The count rates are provided as output data of a so-called detector channel, which is assigned to a detection unit in each case. In the case of quantum counting detectors or photon counting detectors having a plurality of energy thresholds, each detector channel usually generates a set of count rates per projection on the basis of the respective detection signal of the detection unit. The set of count rates in this case may comprise count rates for a plurality of different energy threshold values, in particular energy threshold values which are checked concurrently. The energy threshold values, and the number of energy thresholds to which an energy threshold value is assigned in each case, are usually preset as signal analysis parameters for capturing the projection.
The application of such photon counting detectors in the context of clinical computer tomography allows spectral imaging to take place using polychromatic radiation sources in typically 2 to 4 spectral ranges. The previously cited energy threshold values to be checked correspond to these ranges. Such dual-energy or multi-energy CT methods allow the identification and quantification of different materials such as e.g. iodine and bones in the patient.
In addition to this, it is possible using photon counting detectors to achieve a considerably higher detector resolution, which is approximately two to five times the resolution of conventional CT detectors.
Use of the two advantages of a quantum counting detector, i.e. the spectral imaging and the higher resolution, implies a huge increase in the volume of data to be processed. Said increase relates to both the volume of raw data resulting from an image recording and the subsequent processing of the raw data to produce image data. Due to the transfer of the spectral information, the data rate in the case of 2 to 4 channels increases by a factor of 2 to 4. As a result of increasing the spatial resolution by a factor of 2 to 5, the data rate increases by a factor of 22 to 52. Therefore the volume of data under the cited conditions can increase by a factor of 8 to 100. In addition to this, the so-called frame rate, i.e. the frequency with which the individual image recordings are recorded in the case of a CT image recording from different directions, must be adapted to the higher spatial resolution (i.e. increased) in order that the higher resolution can be used, thereby further increasing the volume of data.
One possibility for limiting the volumes of data during spectral CT imaging is to reduce the detector resolution via so-called pixel fusing, i.e. combining of a plurality of pixels. Alternatively, it is also possible to limit the number of spectral ranges in order to reduce the volumes of data transferred. A further possibility for limiting the volumes of data is to reduce the zoning of a detector. However, the cited measures require the user to decide in advance which properties are particularly important for the imaging. For example, it must be decided how important a maximum resolution or the availability of spectral information or the size of the image region to be depicted are to the examination. Settings must be made accordingly in advance of the image recording, in order to obtain the desired imaging parameters. However, optimization of one of the cited properties is made at the expense of the others and vice versa.